This invention relates to an x-ray sensor having a photoelectric sensor array, particularly a semiconductor array. The invention has particular application to dental intra-oral radiography, although the invention is not so limited.
Charge coupled device (CCD) x-ray imaging sensors with, for example, 912×1368 pixels and an image area of 20 mm×30 mm, are known for digital intra-oral dental applications, and permit a lower x-ray dose than photographic systems. They also produce an image without the delay necessary for photographic development, avoid the storage and use of photographic developing chemicals and facilitate digital archiving of images.
Known dental intra-oral x-ray sensors include an inflexible, sealed, planar, packaged, ccd array for insertion in a mouth to be examined, the package being typically 25 mm×39.5 mm×5.7 mm excluding a cable connection. In use the sensor is typically accommodated in a holder attached to an x-ray source to maintain the sensor at a predetermined distance from, and at a predetermined orientation to, the x-ray source.
Intra-oral dental x-ray imaging sensors have a semiconductor imager die with a close-coupled scintillator layer. In a known x-ray area array sensor, a thin scintillator fluorescent layer is deposited directly onto a CCD or CMOS imager die. Herein, unless the context demands otherwise, fluorescence is to be understood to be radiation-induced luminescence which occurs only during irradiation and phosphorescence is to be understood to be radiation-induced luminescence which persists, and decays, after irradiation. X-ray photons interact with the fluorescent scintillator to stimulate emission of visible wavelength light during irradiation, which is detected and read out from the imager die as an electronic signal. However, some of the x-ray photons pass through the scintillator layer and have, in addition, an unwanted interaction directly with pixels of the imager die. This ‘direct-hit’ interaction occurs relatively infrequently, but when it does occur, the signal generated in a pixel is large. The result is a grainy image and a reduction in an effective signal to noise ratio of the x-ray sensor.
It is clearly desirable to maximize the scintillator interaction and to minimize the direct-hit interaction.
A practical x-ray sensor scintillator requires a balance between x-ray absorption and spatial resolution. It is generally not possible to use a scintillator that is sufficiently thick to stop say 99% of incident x-ray photons. If the scintillator were sufficiently thick, say 500μ, to stop 99% of incident x-rays, there would be considerable absorption and scattering of light photons resulting in little signal and little spatial resolution at the imager die. Hence a thinner scintillator, with less than 99% x-ray absorption, is used, for example 100μ thick.
An x-ray spectrum of a number of photons emitted vs. photon energy from a typical dental x-ray source is a bell-shaped curve with photon energies ranging from 10 keV to 60 keV and a peak number of photons at 32 keV. For a typical 100μ thick scintillator, and x-ray energies between 20 keV and 40 keV, a highest x-ray absorption will be approximately 70%, and a lowest approximately 26%. In other words, between 30% and 74% of x-ray photons incident on the scintillator pass through the scintillator. These x-ray photons give rise to an unwanted secondary interaction if they are stopped by the imaging die, which is typically of silicon. This unwanted interaction produces a relatively large signal charge given by the formula:X-ray photon energy (kev)/3.65i.e. between 2,739 and 16,438 electrons per interaction, i.e. of the order of 100 times the number of imager electrons in a normal scintillator interaction.
In a 20μ active depth in the silicon imaging die approximately 1% of 30 keV photons are stopped.
The silicon imager layer stops around 26 times less photons than the scintillator layer. Transmission loss for x-ray photons through the scintillator is about a factor of 3. Therefore, overall the unwanted x-ray direct-hit interactions are of the order of 100 times less frequent than the wanted scintillator interactions, but when they do occur, they cause of the order of 100 times more signal.
As a consequence, direct hits adversely affect performance of x-ray sensors, and signal to noise ratio of an image including direct-hit interactions is considerably less than one based solely on scintillator interactions. The removal of direct hits produces an image with a higher signal to noise ratio.
As shown in FIG. 1, it is known to employ a fibre-optic plate 102 between a scintillator 101 and an imager die 103 to transmit the light from the scintillator to the imager die while reducing a number of x-ray photons that, having passed through the scintillator, arrive at the imager die. The fibre-optic plate 102, typically 2 mm thick, is bonded between a CCD imager die 103 and the scintillator 101. The 2 mm fibre-optic plate reduces direct hits by of the order of 100 times. However, the fibre-optic plate adds undesirable height or thickness (impacting patient comfort and ease of use, particularly in intra-oral dental applications) and cost, to the x-ray sensor package. Moreover, the fibre-optic plate 102 adds undesirable fixed pattern noise and reduces transmission of light from the scintillator 101 to the imager die 103, increases coupling loss and reduces modulation transfer function (MTF) i.e. resolution.
Similarly to a fibre-optic plate, US-2002/0070365-A1 and U.S. Pat. No. 5,864,146-B disclose the use of an optical grade lead-glass or lead acrylic filter between a scintillator and a CCD imager die. The lead-glass filter absorbs most stray x-rays and prevents them reaching the CCD sensor. US-2002/0070365-A1 and U.S. Pat. No. 5,864,146-B also disclose removal of the CCD imager die from a direct path of an x-ray beam in a linear scanning sensor by use of an arcuate fibre bundle, to reduce a number of x-rays reaching the CCD imager. However, this arrangement may also be expected to suffer from undesirable fixed pattern noise added by the fibre-optic plate and reduced transmission of light from the scintillator to the imager die.
U.S. Pat. No. 5,434,418-B discloses a relatively thick, 200–300μ, complex CsI scintillator layer grown in narrow columns to prevent light spreading in the CsI layer to reduce a probability of x-rays impinging on the silicon imager to less than 0.01%. The semiconductor CCD is formed on a thin 10μ epitaxial layer. Only x-rays absorbed in the epitaxial layer contribute to the image and since silicon is a poor absorber of x-rays of average energy of 35 keV, less than 0.1% incident on the imager die are absorbed in the 10μ top layer.
Use of a storage phosphor from which an image is subsequently read off-line by a semiconductor imager, thereby avoiding exposure of the semiconductor imager to the x-ray beam, is also known.
In particular, EP-1065527-A2 and U.S. Pat. No. 6,504,169-B1 disclose a divalent europium activated caesium halide phosphor screen for storing an image, reading out the image either with a flying spot scanner using a HeNe laser, as illustrated schematically in FIG. 4, or with a scan-head device which reads a line at a time using a row of individual laser diodes. After readout the phosphor screen is erased with an erasing light-source, such as a xenon flash lamp, so that energy remaining in the screen after readout is released to avoid retention of a latent image.
US-2002/0070365-A1 and U.S. Pat. No. 5,864,146-B disclose an optical storage element, such as a photostimulable phosphor, for example barium fluorohalide, which stores an x-ray image and is subsequently illuminated with a second light source, such as a laser or a broadband source, to induce emission of the stored optical energy distribution which is detected by an area detector to provide an image of the subject. Readout may be with a scanned laser beam to emit light which is detected by a photomultiplier tube. Alternatively, a flash-emitting, wide apertured light source 53, as shown schematically in FIG. 5, as opposed to a narrow-apertured laser, may be used, to illuminate the entire imaging area simultaneously to stimulate transmission of a two-dimensional image of the entire x-ray pattern. The image can be recorded using a pixelated CCD in a visible light wavelength sensor 54. A light filter may be used to reject the stimulating light at the sensor 54 and detect only phosphorescent light from the storage phosphor. However, even with the use of filters, problems may be encountered because a portion of the excitation pulse can pass through the filters and on to the CCD array of the light sensor 54. Interference can be avoided by using a flash duration much shorter than that of the stimulated phosphorescent pulse. By time gating, the CCD array can be activated for a controlled interval of time such that the array registers phosphorescence only for a limited time matching duration of the stimulated phosphorescent pulse. Accordingly, interference from the excitation pulse can be minimized by using an excitation pulse duration much shorter than a typical one to five microsecond duration of the stimulated phosphorescent pulse. Because the excitation pulse precedes the phosphorescent pulse, a sufficiently brief excitation pulse is exhausted before the peak of the stimulated phosphorescent pulse is emitted. Time gating can be used to read the CCD in a binned mode at very rapid frame rates and an image frame frozen immediately after decay of the excitation pulse. The signal detected during fast successive framing is discarded and the image signal of the freeze frame which contains the phosphorescent image is retained. Similarly, where a scanning laser beam is used instead of a flashlight, the laser beam may be pulsed and gated with an optical phosphorescence detector to avoid the use of a filter and thereby increase a luminescent signal.
A disadvantage of these off-line reading arrangements is a delay in acquiring images while the storage phosphor is removed from the irradiation site and read and subsequently erased before re-use.
WO 96/16510 appears to disclose reading out of an image sensor after incident light has ceased. In particular, the disclosure relates to an image pick-up apparatus for picking up a series of images at a high image rate. Image read-out commences only after light intensity of the image has nearly vanished, so that hardly any light, e.g. from afterglow of a phosphor layer on an exit window of an image intensifier, is incident on the image sensor during readout. Thus no mechanical shutter is required because the image sensor is not read out as long as the image sensor is being illuminated. Thus the image is integrated throughout x-ray exposure and possibly during some of the afterglow and the image is read after the end of the afterglow.